This invention relates generally to magnetic resonance imaging (MRI) systems and, more particularly, to radio-frequency (RF) coils in such systems.
Magnetic Resonance Imaging (MRI) utilizes hydrogen nuclear spins of the water molecules in the human body, which are polarized by a strong, uniform, static magnetic field of a magnet (typically denoted as B0—the main magnetic field in MRI physics). The magnetically polarized nuclear spins generate magnetic moments in the human body. The magnetic moments point in the direction of the main magnetic field in a steady state and produce no useful information if they are not disturbed by any excitation.
The generation of nuclear magnetic resonance (NMR) signals for MRI data acquisition is accomplished by exciting the magnetic moments with a uniform radio-frequency (RF) magnetic field (typically referred to as the B1 field or the excitation field). The B1 field is produced in the imaging region of interest by an RF transmit coil that is driven by a computer-controlled RF transmitter with a power amplifier. During excitation, the nuclear spin system absorbs magnetic energy, and it's magnetic moments precess around the direction of the main magnetic field. After excitation, the precessing magnetic moments will go through a process of free induction decay (FID), releasing their absorbed energy and returning to a steady state. During FID, NMR signals are detected by the use of a receive RF coil that is placed in the vicinity of the excited volume of the human body. The NMR signal is the secondary electrical voltage (or current) in the receive RF coil that has been induced by the precessing magnetic moments of the human tissue. The receive RF coil can be either the transmit coil itself or an independent receive-only RF coil. The NMR signal is used for producing MR images by using additional pulsed magnetic gradient fields that are generated by gradient coils integrated inside the main magnet system. The gradient fields are used to spatially encode the signals and selectively excite a specific volume of the human body. There are usually three sets of gradient coils in a standard MRI system that generate magnetic fields in the same direction of the main magnetic field and varying linearly in the imaging volume.
In MRI, it is desirable for the excitation and reception to be spatially uniform in the imaging volume for better image uniformity. In a standard MRI system, the best excitation field homogeneity is usually obtained by using a “whole-body” volume RF coil for transmission. The “whole-body” transmit coil is the largest RF coil in the system. A large coil, however, produces lower signal-to-noise ratio (SNR or S/N) if it is also used for reception, mainly because of its greater distance from the signal-generating tissues being imaged. Because a high signal-to-noise ratio is the most desirable in MRI, special-purpose coils are used for reception to enhance the S/N ratio from the volume of interest.
In practice, it is desirable for a well-designed specialty RF coil to have the following functional properties: high S/N ratio, good uniformity, high unloaded quality factor (Q) of the resonance circuit, and high ratio of the unloaded to loaded Q factors. In addition, the coil device must be mechanically designed to facilitate patient handling and comfort, and to provide a protective barrier between the patient and the RF electronics. Another way to increase the SNR is by quadrature reception. In this method, NMR signals are detected in two orthogonal directions, which are in the transverse plane or perpendicular to the main magnetic field. The two signals are detected by two independent individual coils that cover the same volume of interest. With quadrature reception, the SNR can be increased, for example, by up to √{square root over (2)}, over that of the individual linear coils.
A sensitivity encoding (SENSE) technique allows for reducing imaging time by increasing imaging speed. Using the SENSE technique, the spatial sensitivity information in the imaging space (i.e., the real space) provided by the coil elements of a multiple-coil array system can be used to substitute for the information provided by the encoding gradient in the k-space. By skipping some k-space lines, therefore saving imaging time, and using the spatial sensitivity information provided by each of the coil elements, an artifact-free full field of view (FOV) image may still be reconstructed. For example, by eliminating two-thirds of the k-space lines (i.e., by tripling the distance between two adjacent k-space lines), the imaging time can be reduced by about two-thirds. Tripling the distance between two adjacent k-space lines results in a reduction of the FOV in the imaging space to one-third of its original full FOV size. Therefore, the image intensity of each pixel inside the reduced FOV image will be the superposition of the image intensity of three pixels at three different locations in the full FOV image. Knowing the spatial sensitivity profile of each coil element of a multiple-coil array system (at least three coil elements are needed in this case) in the full FOV image and how the reduced FOV image is formed, the superimposed intensities can be separated for each pixel inside the reduced FOV image by solving a set of linear equations. Transferring the separated intensities of the three pixels back to their original locations and performing the same procedures for all the pixels inside the reduced FOV image, the original full FOV image can be reconstructed.
In MRI, a torso RF coil is typically used to image the human torso region, for example, from the top of the liver to the iliac crest or the pelvic region and from the iliac crest to the pubic symphysis. Abdominal and pelvic imaging requires a torso coil to be able to provide good image uniformity in the axial direction (i.e., the transverse direction) as well as good SNR. Non-uniform images caused by the inhomogeneous signal sensitivity profile of a RF coil can lead to misdiagnosis of patients, for example, a high signal in the anterior abdomen region may be mistaken for an enhancing peritoneal tumor. A torso coil is often also used for cardiac imaging. For torso and cardiac SENSE imaging, the sensitivity encoding needs to be in both the left-right and anterior-posterior directions (i.e., the x and y directions). The capability of performing SENSE imaging in the superior-inferior direction (i.e., the head-feet or the z-direction) is also desirable.
Known coil arrangements, and specifically, a birdcage transmit and receive “whole-body” coil of many MRI scanners, can be used to image a patient's abdomen and pelvis with good image homogeneity. However, a major disadvantage of using a “whole-body” coil as a receive coil is that the SNR is too low. The low SNR of a “whole-body” coil is caused by a low filling factor and also by the noise/unwanted signals from the tissue outside the region-of-interest (ROI). The filling factor of a RF coil is determined by the ratio of the volume of the sample (e.g., a human patient's body) being imaged to the effective imaging volume of the coil. The closer the filling factor of a RF coil to unity the better SNR of the coil. Usually, a “whole-body” coil has an effective imaging volume much larger than the volume of the body portion of a patient being imaged. Thus, a “whole-body” coil typically covers a much larger FOV (e.g., about 48 cm) than the body portion of interest to be imaged (e.g., 30 cm for the torso imaging). This causes the “whole-body” coil to couple to more noise and unwanted signals from the tissue outside the ROI and results in a lower SNR. In addition, a “whole-body” coil cannot be used for SENSE imaging.
Known array coils also allow imaging of a large field-of-view (FOV) while maintaining the SNR characteristic of a small and conformal coil. For example, a four-element “C-shaped” adjustable volume array coil is known and that improves the SNR for volume imaging. The mechanical housing of the “C-shaped” volume array coil is divided into two parts: anterior and posterior. Electrically, the “C-shaped” volume array coil consists of four loop coils: three loop coils in the anterior housing and one loop coil in the posterior housing. Each loop coil is critically coupled to its adjacent coil or coils to minimize the inductive coupling between the two adjacent coils and hence to reduce the noise correlation caused by the “cross-talk” between them. However, the four-element “C-shaped” volume array coil cannot provide uniform coverage over the entire axial direction of the torso (i.e., the cross-section of the torso) because it covers only about one half (i.e., the body portion inside the “c-shaped” coil) of the area of the torso cross-section.
It is also known that the direction of the magnetic field generated by a butterfly coil (or saddle coil) can be perpendicular to that generated by a loop coil. Thus, by using a pair of butterfly and loop coils, quadrature detection of a magnetic resonant signal can be achieved. Coil quadrature RF coil systems using this arrangement, including quadrature RF coil systems for neck/c-spine imaging and peripheral vascular imaging are known. The neck/c-spine RF coils typically include two quadrature coil pairs that are placed on the anterior and posterior of the imaging volume (e.g., the neck), respectively. Each of the quadrature coil pairs is formed by a loop coil and a split loop coil and is symmetric about the middle line of the coil. In known peripheral vascular RF coils, three butterfly-loop quadrature pairs are provided. Each of the butterfly-loop quadrature pair includes a large butterfly coil and smaller loop coil positioned at the middle of the butterfly coil. The loop coils are placed under the patient and the flexible wings of the butterfly coils are wrapped around the patient. Because the flexible RF coil system is wrapped around the patient, its filling factor is optimized (i.e., close to unity).
However, the quadrature neck/c-spine coil cannot perform effective SENSE imaging for the entire volume-of-interest (VOI), but only for the middle of the VOI. Further, each of the coil elements of this neck/c-spine coil has left-right symmetry (i.e., symmetric about the middle line of the coil) and also cover more than one-half of the VOI such that the distinctiveness of the complex sensitivity of the coil elements is not sufficient (except for the middle of the VOI) to perform SENSE imaging, particularly in the left-right direction. The peripheral vascular flexible RF coil system provides a much higher signal at the posterior region of the torso than that at the anterior region because the main coil section (i.e., the loop coil and about one half the butterfly coil) is under the patient and only the wings of the large butterfly coil are wrapped around the patient to cover the anterior region. Thus, the signal homogeneity of this flexible coil system in the axial direction is not acceptable for body imaging. Therefore, use of the flexible coil system is limited, for example, used as a vascular coil. In addition, the coil elements of this flexible coil system also only have left-right symmetry and do not distribute in the anterior-posterior direction. Thus, the flexible coil system cannot be used to perform SENSE imaging in both the left-right and anterior-posterior directions.
Further, these know coil arrangements (e.g., four-element ‘C-shaped’ coil, neck/c-spine coil, and peripheral vascular coil) are not dedicated SENSE coils. When used for SENSE imaging, these coils will generate higher geometrical noise (i.e., higher g-factor). Some known RF array coils are optimized for SENSE imaging of the torso and cardiac. For example, a known body coil includes four elliptic shaped loop elements: two on the flexible top and the other two at the bottom. There is no overlap between the adjacent coil elements. The “cross-talk” among the coil elements is minimized by using high input impedance preamplifiers. A known cardiac coil includes four rectangular coil elements, two on the top and the other two at the bottom, and two circular coil elements placed at the left and right, respectively. The two circular lateral coils are also tilted, for example, by 10°, for optimizing performance of SENSE imaging. High input impedance preamplifiers are used to reduce the inductive coupling among the coil elements having no overlap between adjacent coil elements. These optimized torso-SENSE and cardiac-SENSE coils, when being used for SENSE imaging, can provide a much lower geometrical noise than do the conventional torso and cardiac array coils.
However, the torso-SENSE coil and cardiac-SENSE coil are two-dimensional (2D) SENSE coils because the elements of the coils are arranged in the x and y directions. Thus, these coils can only perform SENSE imaging in the left-right direction (x-direction) and the anterior-posterior direction (y-direction) but not in the head-feet direction (z-direction). In order to perform SENSE imaging in the z-direction, coil elements also have to be arranged in the z-direction. Coils, including torso-SENSE imaging coils are also known for sensitivity encoding in all three directions, for example, a torso array coil with eight QD-surface coils for parallel imaging and a cardiac-SENSE imaging array. The three-dimensional (3D) torso-SENSE coil includes eight quadrature coil pairs: four for the anterior section and the other four for the posterior section. Each quadrature coil pair is formed by one rectangular loop coil and one 8-figure coil. Two quadrature coil pairs are arranged in all the x, y, and z directions, which allows the coil to perform SENSE imaging in the all three directions. This requires an eight channel imaging system for operation. The 3D cardiac-SENSE coil includes eight linear loop coils with the inductive coupling between adjacent coil elements distributed in the x-direction minimized by using transformers.
However, the torso-SENSE coil and the cardiac-SENSE coil do not operate satisfactorily as conventional RF coils because they have higher intrinsic noise (or lower intrinsic SNR). To achieve lower geometrical noise for SENSE imaging, a SENSE coil usually does not overlap the adjacent coil elements to critically decouple from each other, but uses high input (or low input) impedance preamplifiers to reduce the inductive coupling among the coil elements. This often results in insufficient isolation among the coil elements and hence higher intrinsic noise or lower intrinsic SNR.
Further, when used as conventional coils, SENSE coils cause image inhomogeneity. To achieve a higher reduction factor for SENSE imaging, the distinctiveness of the complex sensitivity profile of each of the coil elements of a SENSE coil is important. The distinctiveness of the complex sensitivity profile of each coil element of a SENSE coil is usually achieved by using smaller coil elements because of the strong local sensitivity profile. When the 3D torso-SENSE coil and the 3D cardiac-SENSE coil are used as conventional coils (e.g., for conventional imaging without intensity correction) image non-uniformity results. For these 3D SENSE coils, the anterior section needs to be far enough away from the posterior section to enable the coil elements of the two sections to isolate from each other. This can result in shading in the middle of the axial images obtained using these SENSE coils.
Thus, these known coil arrangements are configured such that limited discrimination between field patterns from the separate coil elements is provided. Therefore, when performing 3D parallel imaging, a separate channel is required for receiving signals from each of the coil elements, thus limiting the types of MRI systems capable of performing the 3D parallel imaging.